Method for dynamic stabilization of PET detector gains

ABSTRACT

A method and apparatus for calibrating PET PMTs includes conducting a calibration procedure and determining whether a number of photons absorbed by a corresponding crystal exceeds a count threshold. If the threshold is exceeded the steps of conducting a calibration procedure and determining whether a number of photons absorbed by a corresponding crystal exceeds a count threshold would be repeated and if the threshold is not exceeded then the calibration procedure would be ended.

BACKGROUND OF THE INVENTION

The present invention relates to PET scanners generally and specificallyto a method and apparatus for adjusting PMT gains to compensate fordrift due to various operating phenomenon.

Positrons are positively charged electrons which are emitted byradionuclides which have been prepared using a cyclotron or otherdevice. The radionuclides most often employed in diagnostic imaging arefluorine-18, carbon-11, nitrogen-13 and oxygen-15. Radionuclides areemployed as radioactive tracers called “radiopharmaceuticals” byincorporating them into substances such as glucose or carbon dioxide.One common use for radiopharmaceuticals is in the medical imaging field.

To use a radiopharmaceutical in imaging, the radiopharmaceutical isinjected into a patient and accumulates in an organ, vessel or the like,which is to be imaged. It is known that specific radiopharmaceuticalsbecome concentrated within certain organs or, in the case of a vessel,that specific radiopharmeceuticals will not be absorbed by a vesselwall. The process of concentrating often involves processes such asglucose metabolism, fatty acid metabolism and protein synthesis.Hereinafter, in the interest of simplifying this explanation, an organto be imaged will be referred to generally as an “organ of interest” andprior art and the invention will be described with respect to ahypothetical organ of interest.

After a radiopharmaceutical becomes concentrated within an organ ofinterest and while the radionuclides decay, the radionuclides emitpositrons. Each positron travels a very short distance before itencounters an electron and, when the positron encounters an electron,the positron is annihilated and converted into two photons, or gammarays. This annihilation event is characterized by two features which arepertinent to medical imaging and particularly to medical imaging usingphoton emission tomography (PET). First, each gamma ray has an energy ofessentially 511 keV upon annihilation. Second, the two gamma rays aredirected in substantially opposite directions.

In PET imaging, if the general locations of annihilations can beidentified in three dimensions, the shape of an organ of interest can bereconstructed for observation. To detect annihilation locations, a PETscanner is employed. An exemplary PET scanner includes one or more ringsof detector modules and a processor which, among other things, includescoincidence detection circuitry. The detector modules are arranged aboutan imaging area. An exemplary detector module includes six adjacentdetector blocks. An exemplary detector block includes an array of36bismuth germinate (BGO) scintillation crystals arranged in a 6×6matrix and four photo-multiplier tubes (PMTs) arranged in a 2×2 matrixto the side of the crystal matrix opposite an imaging area.

When a photon impacts a crystal, the crystal generates light which isdetected by the PMTs. The PMT signal intensities are combined togenerate a combined analog signal which is converted into a digitalsignal. For the purposes of this explanation, it will be assumed thatthe digital value, also referred to as a target value, to 511 keV is180. The combined digital signal is compared to a range of values about511 keV. When the combined signal is within the range, an eventdetection pulse (EDP) is generated which is provided to the processorcoincidence circuitry. In addition, acquisition circuits determine whichcrystal within a block absorbed the photon by comparing the relativestrengths of the PMT signals.

The coincidence circuitry identifies essentially simultaneous EDP pairswhich correspond to crystals which are generally on opposite sides ofthe imaging area. Thus, a simultaneous pulse pair indicates that anannihilation has occurred somewhere on a straight line between anassociated pair of crystals. Over an acquisition period of a fewminutes, millions of annihilations are recorded, each annihilationassociated with a unique crystal pair. After an acquisition period,recorded annihilation data is used via any of several different wellknown procedures to construct a three dimensional image of the organ ofinterest.

While operation of a PET detector is relatively simple in theory,unfortunately, despite efforts to manufacture components that operate inan ideal fashion, there is an appreciable variation in how similardetector components respond to identical stimuli. For example, given adetector block including 36 crystals and four PMTs and given the samestimuli, crystals that are positioned proximate the center of the PMTarray will typically generate a higher energy value than edge or cornercrystals (i.e., crystals that are positioned along the edge of the arrayor at the corner of the array). This disparate and position dependentenergy spectrum occurs because, typically, some of the light generatedby an edge or corner crystal is not detected by the PMTs in a singleblock.

As one other example, even within a single crystal, impacting photonsmay not generate the same PMT output for various reasons. For instance,some photons are completely absorbed by a crystal while others are not.Completely absorbed photons generate light corresponding to 511 keVwhile partially absorbed photons generate less than the 511 keV. Asanother instance, first and second photons may be partially absorbedessentially simultaneously by first and second crystals in the sameblock. While each photon would be identified if they had been absorbedconsecutively, upon simultaneous absorption, the combined energy mayerroneously be attributed to a single absorbed photon. In this casedetection circuitry may erroneously identify a third crystal between thefirst and second crystals as the detecting or absorbing crystal.

Thus, while each detected photon should ideally generate a signal havingan energy level of 511 keV, in many cases detected photons generate muchless energy. For this reason, the energy range used to determine if acombined digital PMT signal corresponds to a detected photon typicallyis assigned a relatively low threshold value. For instance, in anexemplary PET system the low end of the energy range may be a digitalvalue of 35 corresponding to approximately 100 keV (i.e., any absorptioneven having an energy greater than 100 keV is assumed to correspond to aphoton).

In addition to the potential errors described above, other sources ofsystem error also occur. For instance, given two PMTs and identicalstimuli (i.e., input light), a first PMT will typically generate aslightly different output signal than the second PMT. Exacerbatingmatters, over time PMT performance has been known to degrade due toaging related changes in structure. Further exacerbating matters, PMTsoften operate differently when exposed to different operatingparameters. For instance, PMT output signals have been known to vary asa function of temperature, ambient magnetic fields and other parametersthat are relatively difficult or expensive to control.

To compensate for PMT construction and operating variances, the PETindustry has developed various commissioning/calibration procedures andassociated hardware and software. Generally, during a calibrationprocedure, a PET source having a known intensity is provided inside thePET imaging area and PMT signals generated thereby are collected. Thecollected PMT signals are compared to expected PMT signals and, wherethere is a difference between the collected and known signals, PMT gainsare adjusted to compensate for the differences.

While calibration techniques like the one described above are useful,unfortunately, most calibration techniques require acquisition ofmassive amounts of data and hence an appreciable amount of time tocomplete. In addition, many calibration techniques include at least somemanual steps that have to be performed by skilled technicians.

Because of the time and skills required to calibrate a PET system, inmany cases, calibration will only be performed when absolutely necessarysuch as after image artifacts begin to appear in generated images. Inother cases calibration is performed routinely whether or not thecalibration is necessary. For instance, in some cases calibration isperformed on a weekly basis. In the case of mobile PET systems (i.e.,truck based systems), the system environment and, in particular, ambientmagnetic fields, may change on a daily basis. In these cases calibrationwill typically be performed on a daily basis.

Thus, in some cases where calibration should be performed, calibrationmay be foregone until a later time while in other cases, wherecalibration is not necessary, a routine calibration procedure may beperformed. In addition, in cases where calibration is only performedwhen a radiologist begins to recognize artifacts, the radiologist isroutinely faced with the question of whether or not to recalibrate.

BRIEF SUMMARY OF THE INVENTION

The above discussed and other drawbacks and deficiencies are overcome oralleviated by a method for calibrating PET detector PMT gains in adetector unit, the method including the steps of conducting acalibration procedure, determining whether a number of photons absorbedby a corresponding crystal exceeds a count threshold, if the thresholdis exceeded, repeating the steps of conducting a calibration procedureand determining whether a number of photons absorbed by a correspondingcrystal exceeds a count threshold, and if the threshold is not exceeded,ending the calibration procedure.

In an alternative embodiment, an apparatus for calibrating PET detectorPMT gains in a detector unit, the apparatus including means forconducting a calibration procedure, means for determining whether anumber of photons absorbed by a corresponding crystal exceeds a countthreshold, means for repeating the calibration procedure if thethreshold is exceeded, and means for ending the calibration procedure ifthe threshold is not exceeded.

In an alternative embodiment, an apparatus for calibrating PET detectorPMT gains in a detector unit, the apparatus including a processor forperforming a pulse sequencing program to perform the steps of conductinga calibration procedure, determining whether a number of photonsabsorbed by a corresponding crystal exceeds a count threshold, whereinif the threshold is exceeded, repeating the steps of conducting acalibration procedure and determining whether a number of photonsabsorbed by a corresponding crystal exceeds a count threshold, andfurther wherein if the threshold is not exceeded, ending the calibrationprocedure.

In an alternative embodiment, a method for improving image performancein PET imaging by stabilizing gain in compensators which are separatelyadjustable so that received digital signals are adjusted to compensatefor PMT degradation, includes calculating PMT signal adjustments withina calibrator and adjusting output of compensators with the PMT signaladjustments, wherein the steps of calculating and adjusting are repeatedduring image acquisition in a PET scanner system until a number ofphotons absorbed by a corresponding crystal does not exceed a countthreshold.

In an alternative embodiment, an imaging system includes a scannersystem, an image reconstruction processor, and ALC circuitry, whereinthe ALC circuitry includes a calibrator for calculating gain adjustmentof compensators within the ALC circuitry and further wherein the imagereconstruction processor includes program signals for defining anexecutable program for repeating calculation of gain adjustment in thecalibrator during image acquisition performed by the scanner system.

The above discussed and other features and advantages of the presentinvention will be appreciated and understood by those skilled in the artfrom the following detailed description and drawings.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a schematic view of a PET system for implementing the presentinvention;

FIG. 2 is a perspective view of a detector unit and associated PMTs.

FIG. 3 is a schematic view of the ALC circuitry of FIG. 1;

FIG. 4 is a flow chart illustrating a commissioning procedure;

FIG. 5 is a flow chart illustrating a calibration procedure;

FIG. 6 is a flow chart illustrating an alternate calibration procedure;

FIG. 7 is a graph illustrating a crystal spectrum generated during acommissioning procedure; and

FIG. 8 is a graph illustrating a combined unit spectrum generated duringa calibration procedure.

DETAILED DESCRIPTION OF THE INVENTION

While various components are described below for carrying out severalinventive methods, it should be appreciated that all of the methodsherein may be performed by any of several different commerciallyavailable and programmable processors.

Referring now to the drawings, wherein like reference characters andsymbols represent corresponding elements and signals throughout theseveral views, and more specifically referring to FIG. 1, the presentinvention will be described in the context of an exemplary PET scannersystem 8. System 8 includes an acquisition system 10, an operator workstation 15, acquisition, locator and coincidence (ALC) circuitry 30 andan image reconstruction processor 40.

System 10 includes a gantry 9 which supports a detector ring assembly 11about a central bore which defines an imaging area 12. A patient table(not illustrated) is positioned in front of gantry 9 and is aligned withimaging area 12. A patient table controller (not shown) moves a tablebed (not shown) into imaging area 12 in response to commands receivedfrom work station 15 through a serial communications link 18.

A gantry controller 17 is mounted within gantry 9 and is responsive tocommands received from operator work station 15 through link 18 tooperate gantry 9. For example, gantry 9 can be tilted away from verticalon command from an operator, can perform a “transmission scan” with acalibrated radio nuclide source to acquire attenuation measurements, canperform a “coincidence timing calibration scan” to acquire correctivedata, or can perform a normal “emission scan” in which positronannihilation events are counted.

As shown best in FIG. 2, assembly 11 is comprised of a large number ofdetector blocks 20. Although not illustrated, detector blocks 20 arearranged in modules, each module including six separate and adjacentdetector blocks 20. A typical assembly 11 includes 56 separate modulessuch that each assembly 11 includes 336 separate detector blocks 20.Each block 20 includes a set of bismuth germinate (BGO) scintillatorcrystals 21 (two separate crystals identified by numerals 180 and 182)arranged in a 6×6 matrix and disposed in front of four photo multipliertubes (PMTs) A, B, C and D which are collectively referred to by numeral22. When a photon impacts a crystal 21, a scintillation event occurs andthe crystal generates light which is directed at PMTs 22. Each PMT 22receives at least some light generated by the scintillation event andproduces an analog signal 23A-23D which arises sharply when ascintillation event occurs and then tails off exponentially with a timeconstant of approximately 300 nanoseconds. The relative magnitudes ofthe analog signals 23A-23D are determined by the position in the 6×6 BGOmatrix at which a scintillation event takes place, and the totalmagnitude of these signals is determined by the energy of an absorbed orpartially absorbed photon which causes the event.

Referring still to FIGS. 1 and 2, a set of acquisition circuitry 25 ismounted within gantry 9 to receive the four signals 23A-23D from eachdetector block 20 in assembly 11. Circuitry 25 provides signals 23A-23Dto ALC circuitry 30 via a data bus 26. Circuitry 30 uses the signals23A-23D to determine the energy E_(i) of a detected event, if the energydetected likely corresponds to a photon, the actual coordinates C_(i) ofa detected event within the block of BGO crystals 21, the time T_(i) ofthe event (i.e. generates a time stamp) and compares event times toidentify coincidence pairs of events that are stored as coincidence datapackets. Each coincidence data packet includes a pair of digital numberswhich precisely identify the addresses of the two BGO crystals 21 thatdetected an associated event. Operation of ALC circuitry 30 is explainedmore in detail below.

Referring again to FIG. 1, processor 40 includes a sorter 34, a memorymodule 43, an array processor 45, an image CPU 42 and a backplane bus 41which conforms to the VME standards and links all other processorcomponents together. The primary purpose of sorter 34 is to generatememory addresses for the coincidence data packets to efficiently storecoincidence data. The set of all projection rays that point in the samedirection and pass through the scanner's FOV is a complete projection,or “view”. A distance R between a particular projection ray and a centerof the FOV locates that projection ray within the FOV. As shown in FIG.1, for example, a positron annihilation (hereinafter an “event”) 50′occurs along a projection ray 51′ which is located in a view at theprojection angle θ and the distance R. Sorter 34 counts all of theevents which occur on this projection ray (R, θ) during an acquisitionperiod by sorting out the coincidence data packets that indicate anevent at the two BGO detector crystals lying on ray 51′.

During data acquisition, the coincidence counts are organized in memory43 as a set of two-dimensional arrays, one for each axial image, andeach having as one of its dimensions the projection angle θ and theother dimension distance R. This θ by R histogram of detected events iscalled a sinogram. Coincidence events occur at random and sorter 34quickly determines the θ and R values from the two crystal addresses ineach coincidence data packet and increments the count of thecorresponding sinogram array element. At the completion of anacquisition period, memory 43 stores the total number of annihilationevents which occurred along each ray (R, θ) in the sinogram.

Image CPU 42 controls bus 41 and links processor 40 to local network 18.Array processor 45 also connects to the bus 41 and operates under thedirection of image CPU 42 to facilitate image reconstruction usinghistogram data from memory module 43. The resulting image array isstored in memory module 43 and may be output by image CPU 42 to operatorwork station 15.

Station 15 includes a CPU 50, a CRT display 51 and a keyboard 52 orother similar input device (e.g., mouse, joystick, voice recognitionmodule, etc.). CPU 50 connects to network 18 and scans key board 52 forinput information. Through the keyboard 52 and associated control panelswitches, an operator can control calibration of system 9, itsconfiguration, and the positioning of a patient table during dataacquisition.

Referring still to FIGS. 1 and 2 and also to FIG. 3, among othercomponents for each block 20, exemplary and simplified ALC circuitry 30includes an analog to digital (AD) converter 95, four compensators 70A,70B, 70C and 70D, an energy, time and crystal identifier 72, coincidencedetection circuitry 90 and event discriminator 91. A separate line(collectively identified by numeral 26) links each of PMTs 22 in a“unit” to the AD converter 95 which converts each of the analog signals23A-23D to a digital signal 23Ad-23Dd. Consistent with the explanationand assumptions above, an analog signal corresponding to 511 keV isconverted to a digital value 180, an analog signal corresponding to 100keV is converted to a digital value 35, etc. Digital values 23Ad-23Ddare provided to compensators 70A-70D, respectively. Each compensator70A-70D is separately adjustable so that the received digital signal(e.g., 23A) may be either increased or decreased to compensate for PMTdegradation or varying operation due to ambient changes.

Compensator outputs are provided to identifier 72 which uses thereceived compensated PMT signals to identify event energy levels E_(i),time T_(i) at which each event occurs and which crystal C_(i) absorbedthe photon that caused the event. Methods and circuitry to perform eachof these tasks are well known in the PET industry and therefore will notbe explained here in detail. For a better understanding of howidentifier 72 operates refer to U.S. Pat. No. 6,232,604 which isentitled Analog Time Adjustment for Coincidence Detection Electronics.

Referring still to FIGS. 1 through 3, the event times T_(i), energiesE_(i) and crystal identifiers C_(i) are provided to event discriminator91 as distinct data packets. Discriminator 91 compares the eventenergies E_(i) to a threshold energy level E_(th) to identify datapackets that likely correspond to valid absorbed photons. In the presentexample, it will be assumed that energy level E_(th) is 100 keVcorresponding to a digital value 35. Thus, discriminator identifies allpackets having digital values between 35 (e.g., 100 keV) and 180 (e.g.,511 keV) and passes times T_(i) and identifiers C_(i) corresponding tothose packets to coincidence detection circuitry 90.

Circuitry 90 compares the times T_(i) in each packet to identifycoincidence events. Where two even times T_(i) are within a small period(i.e., within a “coincidence window”) and if other criteria (e.g.,corresponding crystals C_(i) are separated by the system field of view(FOV)) are met, circuitry 90 identifies the packets as comprising to a“coincidence pair.” Circuit 90 provides coincidence pairs to sorter 34for further image processing as described above.

Hereinafter, in the interest of simplifying this explanation the term“unit” will be used to refer to 25 blocks 20. However, it should beappreciated that the invention contemplates other groupings of blocks.For instance, in some cases a unit may include two blocks 20, fourblocks 20, all blocks 20 that reside in an upper half of detector 11,all blocks that reside in the lower half of detector 11, all blocks 20within detector 11, etc. the smallest unit comprises a single block 20.

Referring to FIG. 3, ALC 30 also includes one or more calibrators 69.During each of a commissioning and a calibration process, calibrator 69receives all energy E_(i) and crystal identifier C_(i) information fromidentifiers 73 (i.e., there is one identifier 72 for each block)corresponding to a single unit. Using the received informationcalibrator 69 calculates PMT signal adjustments used to adjust theoutput of compensators 70A-70D. Thus, in the present example, calibrator69 receives signals from 25 separate blocks 20 (see also FIGS. 1 and 2).

Referring still to FIG. 3, calibrator 69 is used to perform two separateprocesses. A first process, referred to as a commissioning process, isto be performed once or perhaps very seldom (e.g., every month orquarter) to generate information that can be used subsequently duringcalibration processes. The second process, referred to as a calibrationprocess, is meant to be performed routinely. For example, because thecalibration process requires only minimal time to complete, it iscontemplated that a PET system could perform the calibration processbefore every data acquisition procedure and in a manner that isimperceptible to both patient and operator. Alternatively, thecalibration process may be performed repeatedly, even during an imageacquisition procedure. An exemplary commissioning procedure 92 isillustrated in FIG. 4 while an exemplary calibration procedure 114 isillustrated in FIG. 5, and an alternative exemplary calibrationprocedure 140 is illustrated in FIG. 6.

Referring still to FIG. 3, calibrator 69 includes a spectrum generator74, a switch 76, a first peak identifier 78, a gain factor determiner80, a spectrum shifter 82, a spectrum normalizer 84, a second peakidentifier 86, a comparator 88 and an adjuster 89. While shownseparately to simplify this explanation, identifiers 78 and 86 maycomprise a single identifier. Generator 74 and switch 76 are used duringboth the commissioning and calibrating process, identifier 78 anddeterminer 80 are used during the commissioning process and shifter 82,normalizer 84, identifier 86, comparator 88 and adjuster 89 are usedduring the calibration process.

Referring to FIGS. 1, 3 and 4, at block 94 a commissioning photon sourceis provided within imaging area 12 adjacent detector blocks 20. Thesource (not illustrated) directs photons at blocks 20. For each absorbedphoton, identifier 72 provides both the energy E_(i) and the crystalidentifier C_(i) to spectrum generator 74. Throughout a commissioningperiod, generator 74 generates a separate commissioning energy spectrumfor each crystal within the unit. Thus, for instance, referring again toFIG. 2, because a unit includes twenty-five blocks 20 in the presentexample and each block 20 includes 36 crystals (e.g., 180, 182, etc.),generator 74 generates 900 separate crystal spectrums for the exemplaryunit.

Referring now to FIG. 7, an exemplary commissioning spectrum 130 for asingle crystal is illustrated. Referring also to FIG. 2, it will bespectrum 130 corresponding edge crystal 182. Each spectrum 130 plots thenumber of photons absorbed by a corresponding crystal at specific energylevels on a vertical axis against energy level (i.e., E_(i)) on ahorizontal axis. For instance, in exemplary spectrum 130 in FIG. 6,approximately five thousand absorbed photons had energy levelscorresponding to a digital count of 80, approximately twenty thousandabsorbed photons has energy levels corresponding to a count of 120 andapproximately two thousand photons had energy levels corresponding to acount of 148. To generate a spectrum 130 generator 74 simply maintainscounters for each possible digital energy value for each crystal andincrements the appropriate counter when a photon energy level E_(i)matches the level associated with the counter.

Exemplary commissioning spectrum 130 clearly illustrates that theenergies E_(i) attributed to separate absorbed photons vary widely evenwithin a given crystal due to phenomenon described above includingpartial absorption, dual absorption, partial light detection, etc.Clearly there is a peak energy level E_(p) at which a curve through thecount values is at a highest point. In spectrum 130 the peak energylevel occurs at approximately 115.

As indicated above, the digital value attributable to a completelyabsorbed and detected photon is 180 (corresponding to 511 keV). Giventhis assumption, the peak energy level 115 seams to be relatively low asone would expect the peak level to have been approximately 180. As itturns, crystal 182 (see FIG. 2) to which exemplary spectrum 130corresponds is an edge crystal (i.e., a crystal residing along an edgeof a corresponding block 20) which means much of the light generatedthereby is not detected by PMTs. A spectrum corresponding to a morecentrally located crystal (i.e., a crystal near the center of array 21in FIG. 2) would have a peak energy level almost exactly at 180.

Referring again to FIGS. 1, 3 and 4, during the commissioning procedureswitch 76 is closed to identifier 78. At block 98, for each unitcrystal, identifier 78 determines the peak energy level E_(p) of thecorresponding commissioning spectrum 130. Once again, for the spectrum130 in FIG. 6, the peak level is approximately 115. The peak levels areprovided to determiner 80.

Determiner 80 also receives an energy target input E_(t) which indicatesa target energy level for each crystal that corresponds to 511 keV. Thetarget level in the present case is 180 which is provided by a system 8user.

At step 100, determiner 80 combines the target energy level E_(t) witheach of the separate peak energy levels E_(p) for each crystal therebygenerating a separate gain factor G_(f) for each unit crystal. Thiscombining step includes dividing the target level E_(t) by each of thepeak levels E_(p). For instance, in the case of the crystalcorresponding to spectrum 130 in FIG. 6 with a peak level E_(p) of 115,the gain factor G_(f) would be 1.57 (i.e., 180/115=1.57). The gainfactor G_(f) is a factor by which the energies in the commissioningspectrum for the corresponding crystal have to be shifted in order for ashifted peak energy level E_(p) to be equal to the target energy levelE_(t). Thus, in the present example, by multiplying each energy level inspectrum 130 by factor G_(f)=1.57, a compensated spectrum having adesired peak at 180 results.

The gain factors G_(f) are stored for each separate crystal at step 102.During calibration, factors G_(f) are provided to shifter 82.

Referring now to FIGS. 1, 3, 5, and 6, after the commissioning procedurehas been completed and gain factors G_(f) for each crystal stored,calibration process 114 is performed to adjust compensators 70A through70D prior to each data acquisition procedure or on a potentiallyrepeated basis as shown by calibration process 140 within FIG. 6.

At block 104 a radionuclide is provided within a patient for imagingpurposes and the patient is positioned within area 12 adjacent detectorblock 20 and, specifically, adjacent the 25 block unit in the presentexample so that photons are directed at unit crystals during thecalibration process. It has been found that when a patient who has beeninjected with a radionuclide is near an imaging bore, even 0.1millicurie of activity results in a block count rate of approximately200 counts per second. It has also been found that approximately 5000counts are needed in a spectrum to provide desired precision. In thepresent case, where the unit includes 25 blocks, for the unit, 5000counts can be obtained in approximately one second. Thus, in the presentcase, the data acquisition portion of the calibration process would onlyrequire approximately one second.

Referring still to FIGS. 1, 3, 5, and 6 at block 106, during acalibration period, generator 74 receives the energy E_(i) and crystalidentifier C_(i) signals for every crystal within the unit and generatesa separate calibration energy spectrum for each separate crystal. Thecalibration spectrums are similar to commissioning spectrum 130 in FIG.7 plotting counts against energy values E_(i) to form the spectrum. Theprimary difference between the calibration and commissioning spectrumsgenerally is that the count values will be much greater for thecommissioning spectrum than for the calibration spectrum because thecommissioning period (e.g., several minutes) is much longer than thecalibration period (e.g., 1 or 2 seconds). During calibration switch 76is open to identifier 78 and closed to shifter 82.

At block 108, after the calibration spectrums have been generated andstored for each unit crystal, spectrum shifter 82 receives the spectrumsand the crystal specific gain factors G_(f) and multiplies the energylevels in the calibration spectrum by the gain factors G_(f). Forinstance, assume that it has been some time since the commissioningprocedure was performed to generate the gain factors G_(f) and thatcrystal performance has degraded somewhat. In this case, a calibrationspectrum for crystal 182 may have slid such that a peak energy level forthe calibration spectrum is approximately 110 (i.e., the peak has slid 5from the peak level of the commissioning spectrum). Here, themultiplying step 108 would multiply the calibration spectrum for crystal182 by crystal specific gain factor 1.57 thereby shifting the entirecalibration spectrum and generating a shifted spectrum for the crystal.

Thus, upon shifting, the peak energy level for the shifted spectrum willbe similar to value 180. In the present case, the peak level for theshifted spectrum would have a value 172.7. Other spectrum energy levelsare similarly shifted by factor 1.57.

It should be appreciated that, while crystal specific data collectedduring a short calibration acquisition is used to generate a crystalspecific spectrum, because only a small amount of data (e.g., 2-3hundred counts) for the crystal can be collected during the shortcalibration period, the crystal specific spectrum alone is notstatistically very inaccurate.

At block 110 normalizer 84 receives the shifted spectrums from each unitcrystal and combines all of the data corresponding to the shiftedspectrums into a single normalized unit spectrum. An exemplary unitspectrum 132 is illustrated in FIG. 8 where all of the data points fromthe separate crystal spectrum are overlaid onto a single graph and acurve 134 is formed therefrom. In FIG. 8 it can be seen that the peakenergy level of the unit spectrum 132 is approximately 175.

Continuing, at block 112, identifier 86 determines the peak unit energylevel E_(up) for the unit spectrum.

At block 116, comparator 88 compares the peak unit energy value E_(pu)(e.g., 175 in the present example) with the target energy value E_(t)(e.g., 180) and determines the percent difference between the twovalues. For instance, in the present case, where the peak unit valueE_(pu) is 175 and the target value is 180, the difference isapproximately 2.78%.

The difference value is provided to adjuster 89 which, at block 118,adjusts the PMT gains for all of the unit PMTs via compensators 70A-70Dto increase the gains as a function of the difference value. In order toavoid oscillations, adjuster will typically be set to modify the PMTgains by a percentage of the difference value. For example, an exemplarypercentage may be 75% so that, where the difference value is 2.78%, theadjustments would increase gains by 2.09%.

It should be appreciated that, while insufficient counts are collectedon a crystal by crystal basis to provide statistical certainty requiredfor spectral analysis, where counts from many crystals are combined, thenumber of counts is sufficient to facilitate acceptable accuracy despitea short (e.g., one second) calibration acquisition period. Thus,calibration can be performed quickly, relatively accurately and withoutoperator interaction.

Turning now to FIG. 6, a calibration procedure 140 similar to thecalibration procedure 114 is shown. In FIG. 6, however, the calibrationprocedure is preferably initially performed prior to image acquisitionand then may be repeated such as during the image acquisition procedure.Such a continual calibration procedure provides continuous update ofgain to minimize impact on patient imaging efficiency and maximizecapability to maintain stable adjustment over long periods of imageacquisition, which may even last an hour or more. If the sufficientactivity is detected, as determined by step 120, then the calibrationprocedure 140 will be repeated starting at step 106. By “sufficientactivity”, it should be understood that the gain adjustment does requirethe presence of some positron activity, and therefore there should besufficient counts to warrant a recalculation. This makes it impracticalto perform the calibration procedure all day long, but instead isrepeated as long as there is sufficient activity to make the updateworthwhile. A specific count threshold may be set by an operator or maybe preprogrammed within the processor 40. When there are sufficientcounts to perform an update, when the number of counts exceeds the countthreshold, the sampling would occur every few minutes and updating ifthere are sufficient counts observed to calculate a change andsufficient change observed to justify an adjustment. To that end, thealgorithm 140 preferably detects low levels of activity and suspendsupdates at that time, that is, when the counts are less than the countthreshold. It is possible that, even during patient acquisition, ifinsufficient activity is detected, as determined by step 120, then thecalibration procedure 140 will end. By “end”, it should be understoodthat the calibration procedure 140 will again be implemented prior toanother image acquisition procedure.

An assembly incorporating the calibration procedure 140 will include thehardware capability to acquire energy spectra concurrently with patientcoincidence data. Yet another alternative implementation may include thesame hardware as shown while incorporating the ability to switchperiodically and rapidly from a mode of patient acquisition to one ofspectra collection and then back, although repeated and continualadjustment, as deemed necessary by a count threshold, as describedwithin FIG. 6 is preferred.

It should be understood that the methods and apparatuses described aboveare only exemplary and do not limit the scope of the invention, and thatvarious modifications could be made by those skilled in the art thatwould fall under the of the invention.

To apprise the public of the scope of this invention, the followingclaims are made:

1. A method for calibrating PET detector PMT gains in a detector unit,the method comprising the steps of: (a) conducting a calibrationprocedure; (b) determining whether a number of photons absorbed by acorresponding crystal exceeds a count threshold; (c) if the threshold isnot exceeded, repeating steps (a) and (b); and, (d) if the threshold isnot exceeded, ending the calibration procedure wherein the detector unitincludes at least one detector block, where a block includes a twodimensional crystal array including crystal arranged adjacent an imagingarea and a PMT array including a two dimensional array of a PMTsarranged adjacent the crystal array opposite the imaging area, a targetenergy level being associated with the known average energy of a photon,the method further comprising the steps of: providing a calibrationphoton source adjacent the at least one block during a calibrationperiod and, for each unit crystal: (i) obtaining a calibration energyspectrum where the calibration spectrum indicates the number detectedphotons at each of several possible energy levels; and (ii)mathematically combining the calibration spectrum and a crystal specificgain factor to generate a shifted spectrum for the crystal; combiningthe shifted spectrums for all unit crystals to generate a unit spectrum;identifying a peak unit energy level for the unit spectrum where thepeak unit energy level is the energy level at which the greatest numberof photons was detected; comparing the peak unit energy level and thetarget energy level; and based on the difference between the peak unitenergy level and the target energy level, adjusting the PMT gains forthe units PMTs.
 2. The method of claim 1 further including the steps of,prior to the step of providing and during a commissioning procedure:providing a commissioning photon source adjacent the at least one blockand during a commissioning period, for each unit crystal: (i) obtaininga commissioning energy spectrum where the commissioning spectrumindicates the number of detected photons at each of several possibleenergy levels; (ii) identifying a peak energy level for thecommissioning spectrum; and (iii) mathematically combining the targetenergy level and the peak energy level to generate the crystal specificgain factor.
 3. The method of claim 2 wherein the step of providing thecalibration source includes the steps of providing a radionuclide withina patient and positioning the patient adjacent the crystal array.
 4. Themethod of claim 2 wherein the step of mathematically combining togenerate the gain factor includes the step of dividing the target energylevel by the peak energy level.
 5. The method of claim 4 wherein thestep of mathematically combining to generate the shifted energy spectrumincludes the step of multiplying each energy level within the spectrumby the crystal gain factor thereby shifting each of the energy levelcounts.
 6. The method of claim 5 wherein the step of comparing includesthe step of determining the percentage difference between the peak unitenergy level and the target energy level and wherein the step ofadjusting includes the step of adjusting the gain of each of the PMTs ina manner calculated to modify the gains by the percentage difference. 7.The method of claim 1 wherein the unit is a first unit and the PETdetector includes at least a second unit and wherein the process isperformed simultaneously for each detector unit to adjust unit PMTgains.
 8. The method of claim 7 wherein the first unit is positionedabove the second unit in the detector.
 9. The method of claim 1 whereinthe calibration period is between one half second and ten seconds, andis repeated until the threshold is not exceeded.
 10. A method forcalibrating PET detector PMT gains in a detector unit including at leastone detector block, where a block includes a two dimensional crystalarray including crystals arranged adjacent an imaging area and a PMTarray including a two dimensional array of PMTs arranged adjacent thecrystal array opposite the imaging area, a target energy level beingassociated with the known average energy of photon, the methodcomprising the steps of: (a) conducting a calibration procedure; (b)determining whether a number of photons absorbed by a correspondingcrystal exceeds a count threshold; (c) if the threshold is exceed,repeating steps (a) and (b); and, (d) if the threshold is not exceeded,ending the calibration procedure; the method further comprising thesteps of: (A) providing a commissioning photon source adjacent the atleast one block and, during a commissioning period, for each unitcrystal: (i) obtaining a commissioning energy spectrum where thecommissioning spectrum indicates the number of detected photons at eachof several possible energy levels; (ii) identifying a peak energy levelfor the commissioning spectrum at which the greatest number of photonswas detected; and (iii) mathematically combining the target energy leveland the peak energy level to generate a crystal specific gain factor;(B) providing a radionuclide within a patient and positioning thepatient adjacent the crystal array and, during a calibration period, foreach unit crystal: (i) obtaining a calibration energy spectrum where thecalibration spectrum indicates the number of detected photons at each ofseveral possible energy levels; and (ii) mathematically combining thecalibration spectrum and the crystal specific gain factor to generate ashifted spectrum for the crystal; combining the shifted spectrums forall unit crystals to generate a unit spectrum; identifying a peak unitenergy level for the unit spectrum; comparing the peak unit energy leveland the target energy level; and based on the difference between thepeak unit energy level and the target energy level, adjusting the PMTgains for the unit PMTs.
 11. The method of claim 10 wherein the step ofmathematically combining to generate the gain factor includes the stepof dividing the target energy level by the peak energy level.
 12. Themethod of claim 11 wherein the step of mathematically combining togenerate the shifted energy spectrum includes the step of multiplyingeach energy level within the spectrum by the crystal gain factor therebyshifting each of the energy level counts.
 13. An apparatus forcalibrating PET detector PMT gains in a detector unit, the apparatuscomprising: means for conducting a calibration procedure; means fordetermining whether a number of photons absorbed by a correspondingcrystal exceeds a mount threshold; means for repeating the calibrationprocedure if the threshold is exceeded; and means for ending thecalibration procedure of the threshold is not exceeded; the apparatusfurther including at least one detector block, where a block includes atwo dimensional crystal array including crystals arranged adjacent animaging area and a PMT array including a two dimensional array of PMTsarranged adjacent the crystal array opposite the imaging area, a targetenergy level being associated with the known average energy of a photon,the apparatus further comprising: means for providing a calibrationphoton source adjacent the at least one block during a calibrationperiod and, for each unit crystal: (i) means for obtaining a calibrationenergy spectrum where the calibration spectrum indicates the number ofdetected photons at each of several possible energy levels; and (ii)means for mathematically combining the calibration spectrum and acrystal specific gain factor to generate a shifted spectrum for thecrystal; means for combining the shifted spectrums for all unit crystalsto generate a unit spectrum; means for identifying a peak unit energylevel for the unit spectrum where the peak unit energy level is theenergy level at which the greatest number of photons was detected; meansfor comparing the peak unit energy level and the target energy level;and means for, based on the difference between the peak unit energylevel and the target energy level, adjusting the PMT gains for the unitPMTs.
 14. The apparatus of claim 13 further including means forperforming a commissioning procedure including: means for providing acommissioning photon source adjacent the at least one block and during acommissioning period, for each unit crystal: (i) means for obtaining acommissioning energy spectrum where the commissioning spectrum indicatesthe number of detected photons at each of several possible energylevels; (ii) means for identifying a peak energy level for thecommissioning spectrum; and (iii) means for mathematically combining thetarget energy level and the peak energy level to generate the crystalspecific gain factor.
 15. The apparatus of claim 14 wherein the meansfor mathematically combining to generate the gain factor includes meansfor dividing the target energy level by the peak energy level.
 16. Theapparatus of claim 15 wherein the means for mathematically combining togenerate the shifted energy spectrum includes means for multiplying eachenergy level within the spectrum by the crystal gain factor therebyshifting each of the energy level counts.
 17. An apparatus forcalibrating PET detector PMT gains in a detector unit, wherein thedetector unit includes at least one detector block, where a blockincludes a two dimensional crystal array including crystal arrangedadjacent an imaging area and a PMT array including a two dimensionalarray of PMTs arranged adjacent the crystal array opposite the imagingarea, a target energy level being associated with the known averageenergy of a photon, the apparatus comprising: a processor for performinga pulse sequencing program to perform the steps of: (a) conducting acalibration procedure; (b) determining whether a number of photonsabsorbed by a corresponding crystal exceeds a count threshold; (c) ifthe threshold is exceeded, repeating steps (a) and (b); (d) if thethreshold is not exceeded, ending the calibration procedure; wherein thepulse sequencing program further causes the processor to perform thesteps of: providing a calibration photon source adjacent the at leastone block during a calibration period and, for each unit crystal: (i)obtaining a calibration energy spectrum where the calibration spectrumindicates the number of detected photons at each of several possibleenergy levels; and (ii) mathematically combining the calibrationspectrum and a crystal specific gain factor to generate a shiftedspectrum for the crystal; combining the shifted spectrums for all unitcrystals to generate a unit spectrum; identifying a peak unit energylevel for the spectrum where the peak unit energy level is the energylevel at which the greatest number of photons was detected; comparingthe peak unit energy level and the target energy level; and based on thedifference between the peak unit energy level and the target energylevel, adjusting the PMT gains for the unit PMTs.
 18. The apparatus ofclaim 17 wherein the pulse sequencing program further causes theprocessor to perform the steps of, prior to the step of providing andduring a commissioning procedure: providing a commissioning photonsource adjacent the at least one block and during a commissioningperiod, for each unit crystal: (i) obtaining a commissioning energyspectrum where the commissioning spectrum indicates the number ofdetected photons at each of several possible energy levels; (ii)identifying a peak energy level for the commissioning spectrum; and(iii) mathematically combining the target energy level and the peakenergy level to generate the crystal specific gain factor.
 19. Theapparatus of claim 18 wherein, to perform the step of mathematicallycombining to generate the gain factor the program causes the processorto perform the step of dividing the target energy level by the peakenergy level.
 20. The apparatus of claim 19 wherein, to perform the stepof mathematically combining to generate the shifted energy spectrum theprogram causes the processor to perform the step of multiplying eachenergy level within the spectrum by the crystal gain factor therebyshifting each of the energy level counts.
 21. A method for improvingimage performance in PET imaging by stabilizing gain in compensatorswhich are separately adjustable so that received digital signals areadjusted to compensate for PMT degradation, the method comprising: (a)calculating PMT signal adjustments within a calibrator; and, (b)adjusting output of compensators with the PMT signal adjustments;wherein steps (a) and (b) are repeated during image acquisition in a PETscanner system until a number of photons absorbed by a correspondingcrystal does not exceed a count threshold; wherein the detector unitincludes at least one detector block, where a block includes a twodimensional crystal array including crystal arranged adjacent an imagingarea and a PMT array including a two dimensional array of PMTs arrangedadjacent the crystal array opposite the imaging area, a target energylevel being associated with the known average energy of a photon, themethod further comprising the steps of: providing a calibration photonsource adjacent the at least one block during a calibration period and,for each unit crystal: (i) obtaining a calibration energy spectrum wherethe calibration spectrum indicates the number of detected photons ateach of several possible energy levels; and (ii) mathematicallycombining the calibration spectrum and a crystal specific gain factor togenerate a shifted spectrum for the crystal; combining the shiftedspectrums for all unit crystals to generate a unit spectrum; identifyinga peak unit energy level for the unit spectrum where the peak unitenergy level is the energy level at which the greatest number of photonswas detected; comparing the peak unit energy level and the target energylevel; and based on the difference between the peak unit energy leveland the target energy level, adjusting the PMT gains for the unit PMTs.22. The method of claim 21 wherein steps (a) and (b) are performedsimultaneously with image acquisition.
 23. An imaging system comprising:a scanner system; an apparatus for calibrating PET detector PMT gains ina detector unit, wherein the detector unit includes at least onedetector block, where a block includes a two dimensional crystal arrayincluding crystal arranged adjacent an imaging area and a PMT arrayincluding a two dimensional array of PMTs arranged adjacent the crystalarray opposite the imaging area, a target energy level being associatedwith the known average energy of a photon, an image reconstructionprocessor for performing a pulse sequencing program, and, acquisition,locator and coincidence circuitry, wherein the circuitry includes acalibrator for calculating gain adjustment of compensators within thecircuitry and further wherein the image reconstruction processorincludes program signals for defining an executable program forrepeating calculation of gain adjustment in the calibrator during imageacquisition performed by the scanner system; wherein the pulsesequencing program further causes the processor to perform the steps of:providing a calibration photon source adjacent the at least one blockduring a calibration period and, for each unit crystal: (i) obtaining acalibration energy spectrum where the calibration spectrum indicates thenumber of detected photons at each of several possible energy levels;and (ii) mathematically combining the calibration spectrum and a crystalspecific gain factor to generate a shifted spectrum for the crystal;combining the shifted spectrums for all unit crystals to generate a unitspectrum; identifying a peak unit energy level for the unit spectrumwhere the peak unit energy level is the energy at which the greatestnumber of photons was detected; comparing the peak unit energy level andthe target energy level; and based on the difference between the peakunit energy level and the target energy level, adjusting the PMT gainsfor the unit PMTs.